Implantable Biosensor and Methods of Use Thereof

ABSTRACT

Disclosed herein is an analyte sensing device capable of continuously monitoring metabolic levels of a plurality of analytes. The device comprises an external unit, which, for example, could be worn around the wrist like a wristwatch or could be incorporated into a cell phone or PDA device, and an implantable sensor platform that is suitable, for example, for implantation under the skin. The external device and the internal device are in wireless communication. In one embodiment, the external device and the internal device are operationally linked by a feedback system. In one embodiment, the internal device is encapsulated in a biocompatible coating capable of controlling the local tissue environment in order to prevent/minimize inflammation and fibrosis, promote neo-angiogenesis and wound healing and this facilitate device functionality.

CROSS REFERENCE TO RELATED APPLICATION

This application is a continuation application of co-pending U.S.Non-Provisional patent application Ser. No. 11/862,866 filed Sep. 27,2007 and claims priority benefit of the filing date of U.S. ProvisionalApplication Ser. No. 60/827,104 filed Sep. 27, 2006, the contents ofboth of which are hereby incorporated by reference in their entireties.

BACKGROUND

Careful metabolic monitoring and proper treatment can improve control ofmetabolic diseases such as diabetes and obesity. Knowing a patient'smetabolism along with other physiological parameters allows for correctdosing and delivery of medications and nutrients. Improvements inmetabolic measurement technology are essential for better diagnosticsand advances in treatment of metabolic diseases and conditions.Treatment of metabolic diseases and conditions ideally requires frequentand timely monitoring which drives a need for monitors that arenon-invasive, real-time, portable, low cost, and accurate. Metabolicdata are also useful in assessing the physiological homeostaticconditions of patients and healthy subjects in general.

Blood glucose concentration data is extremely useful for the control ofmetabolic diseases such as diabetes and for monitoring the overallmetabolic condition of a human subject. An accurate, real-time,noninvasive method for measurement of blood glucose levels is of greatinterest in the diabetic community. Current technologies involving themeasurement of blood glucose by drawing blood are invasive and oftenlead to poor patient compliance. Measurement by probe involves frequentlancing and may result in problems. An ideal non-invasive blood glucosesensor provides a continuous signal and/or a signal on demand that canbe used to control devices, such as insulin pumps in closed loopfeedback applications.

In recent years, two different types of metabolic internal units havebeen developed: non-invasive and minimally invasive. Non-invasiveoptical internal units depend on light penetration into the skin andspectroscopic measurement of metabolic levels; however, lack of analytespecificity remains a problem for optical internal units. Commerciallyavailable minimally invasive internal units can function only for theshort term (a few days) and require frequent calibration via fingerpricking. These commercially available internal units are eitherincapable of continuous monitoring of metabolic levels or are onlysuitable for use by qualified medical personnel.

Therefore, there exists a need for a minimally invasive or non-invasivemetabolic internal unit suitable for use by the host that allowscontinuous and/or on demand monitoring of metabolic levels of specificanalytes.

SUMMARY

An analyte sensing device comprises an external control unit and animplantable sensor platform in wireless optical communication, whereinthe implantable sensor platform can pass though a 14 gauge or smallerbore needle. This implantable sensor platform comprises a variety offunctional optoelectronic circuit blocks for wireless powering,interactive communication, programmable potentiostats interfacing withvarious electrochemical sensors, mode-selection, signal processing,calibration, analog to digital conversion, amplification, and opticaltransmission. The outer surface of this miniaturized sensor platform iscoated with one or more biocompatible coatings, optionally capable ofreleasing a variety of drugs and tissue response modifiers. The externalcontrol unit comprises optical sources suitable for powering theimplantable sensor platform, along with transmitters and receivers fortransmitting and receiving optical commands to and from the implantablesensor platform. These optical signals are then converted to electricalpulses and processed by a microprocessor located in the external unit.In addition, the external unit is equipped with a miniaturized camera toassist in aligning the various optical components of the external unitwith that of the implantable sensor platform.

DESCRIPTION OF THE FIGURES

FIG. 1. Schematic representation of an embodiment of an implantedbiosensor unit along with an embodiment of an external user interfaceunit comprising mode select, monitoring and calibration functions.

FIG. 2. Schematic representation of an embodiment of a sensor platformshown as stack of three chips encased in a suitable biocompatiblecoating. The sensor platform is compact enough for implantation byneedle and plunger.

FIG. 3. Schematic view of an embodiment of a three sub-chip sensorplatform along with its interface with the external control unit.

FIG. 4. Embodiment of optical and optoelectronic components housedwithin PDA unit and methodology to optically communicate with theimplanted sensor platform.

FIG. 5. Schematic of a programmable potentiostat interfacing with twosensors whose signal is processed by the signal-processing unit. Theoptoelectronic transmitter and receiver interface communicating with theimplanted chip and the modified PDA is also shown.

FIG. 6. Schematic of a PDA communicating wirelessly with an externalunit that is located in the vicinity of the implantable platform.

FIG. 7. Embodiment of a design of sensor-select circuit. This circuitconsists of a optical pulse receiving system, a timer, set of D-FlipFlops, and a logic block. It interfaces calibration and Mux(multiplexer) circuits.

FIG. 8. Figure A showing an ADC signal processor (MOSIS fabricated chip)interfaced with a hybrid potentiostat. (B). Measurements showing digitalsignal changing its pulse characteristics as a function of glucoselevel. (C). Plot of glucose level after converting the pulse frequencychange.

FIG. 9. Photographs showing various components of subchips: solar cells,laser as transmitters, signal processing chips (two generation),three-electrode electrochemical sensors.

FIG. 10. (top) Hybrid approach to integrate sub-chips with robustinterconnects and bonding using vias and bumps. (bottom left)Cross-sectional schematic illustration of bonding between a via and bumpfor sub-chip #1 and sub-chip #2. Top view of interconnect and power padis also shown. (bottom right) Bonding between sub-chip #2 and sub-chip#3 is shown using two approaches for integration.

FIG. 11. Schematic of an embodiment of the three sub-chip design (shownin FIGS. 1, 2 and 3). Here the top surface of sub-chip #1 (4) ishermetically sealed using a transparent glass window (111) sealed to theraised silicon walls (110) via anodic bonding (112) between glass andsilicon.

FIG. 12. Embodiment of a sensor platform having two sub-chips on its topand bottom surfaces. Sub-chip #1 has three pads on either side for powersupply distribution (e.g., V_(dd), V₁, and C for common, shown in blueand are larger in size than the via/bumps pads). The power is suppliedto the sub-chip #2 using metalized vias labeled as V_(VDD) etc. Inaddition, vias are used to connect photodetector PD_(SS)/PD_(M) onsub-chip #1 to sub-chip #2 (having electronics such as sensor select,routing logic/MUX, etc.). Note that the 1.55 μm transmitter is locatedon sub-chip #2 as this wavelength is transparent to Si platform andchips.

FIG. 13. Circuit schematic showing various functions of three sub-chipscondensed into two sub-chips. In this version, the modified PDA unitdirectly communicates with the implanted unit.

FIG. 14. Circuit schematic showing various functions of three sub-chipscondensed into two sub-chips. In this version, the modified PDA unitcommunicates with an external unit located in the vicinity of theimplanted sensor and the communication is via Bluetooth® wirelesstechnology.

FIG. 15. Advanced methodology to integrate two sub-chips into one-waferplatform using wafer bonding technique. The hermetical seal usingglass-Si anodic bonding is also shown for the solar cell/PDs chip.

FIG. 16. Schematic of implantable sensor platform interfacing with adrug dispensing system. The drug delivery may include amicro-electro-mechanical (MEM) components.

FIG. 17. Glucose sensor showing Ag reference electrode and details ofcoatings on Pt working electrode. Schematic representation of modifiedClark amperometric glucose sensor, along with various chemical,electrochemical and diffusion processes associated with its operation.The glucose oxidase (GO_(x)) layer is coated with a semipermeablemembrane to reduce the amount of glucose entering the sensor. AnHRP-modified hydrogel layer is then applied to eliminate outer diffusionof H₂O₂. This is followed by an outer composite hydrogel coating withembedded microspheres at different stages of degradation and TRMrelease.

FIG. 18. Schematic of the self assembly of semipermeable membranecomposed of humic acids and Fe³⁺ on the outer surface of theelectrochemical sensor. As the number of bilayers increases, there is anincrease in the tortuosity for glucose diffusion towards the enzyme.

FIG. 19. One-second pulsed mode operation of sensor. Calibration curveof current response vs. change in glucose concentration. As the amountof glucose in the system is increased, there is a corresponding rise inthe current. The Figure on left indicates the time required for thedevice to stabilize at each glucose concentration. As the concentrationof glucose increases, the device reaches a stable reading faster.

FIG. 20. Current response as a function of increasing semipermeablemembrane thickness.

FIG. 21. Histological evaluation of subcutaneous tissue samples takenfrom the vicinity of hydrogel composites containing PLGA microspheres at3 and 21 days post implantation. The representative sections shown are 3days after implantation (A & B) and 21 days after implantation (C & D).

FIG. 22. schematic of a methodology to hermetically seal sub-chips usingone Si wafer as the carrier with provision to placing sub-chips andinterconnecting them. The hermetical seal using glass-Si anodic bondingis shown on top as well as bottom.

DETAILED DESCRIPTION

Disclosed herein is a device capable of monitoring the metabolic levelsof a plurality of analytes, in a continuous or intermittent (e.g., ondemand) operation. The device comprises an external unit, which, forexample, is worn around the wrist like a wristwatch or carried like aPersonal Digital Assistant (PDA) or a cell phone, and a sensor platformthat is suitable for implantation under the skin or near the surface ofanother portion of a patient's anatomy. The sensor may be implantablevia a needle and similarly removable via a needle, thus avoiding theneed for surgical implantation and removal.

The term “analyte” refers to a substance or chemical constituent in abiological fluid (e.g., blood, interstitial fluid or urine) that can beanalyzed. In one embodiment, the analyte for measurement by the devicesand methods disclosed herein is glucose.

“Biocompatibility” is the ability of a material to perform with anappropriate host response in a specific application. The terms“biocompatible membrane”, “biocompatible layer,” and the like refer to asemipermeable membrane comprised of protective biomaterials. In oneembodiment, a biocompatible membrane is a few microns thickness or moreand is permeable to small-molecule analytes oxygen and glucose, but issubstantially impermeable to biofouling agents (such as proteins) thatcould otherwise gain proximity to and possibly damage the internal unit.This “biocompatible membrane,” or “biocompatible layer,” may alsoprotect the sensor from damage and inconsistency in readings resultingfrom inflammation and fibrous encapsulation. In some embodiments, thebiocompatible membrane comprises pores (e.g., typically fromapproximately 0.1 to approximately 1.0 micron).

An “electrochemical sensor” is a sensor configured to detect thepresence and/or measure the level of an analyte in a sample viaelectrochemical oxidation and reduction reactions on the sensor. Thesereactions are transduced to an electrical signal that can be correlatedto an amount, concentration, or level of an analyte in the sample.

The sensor platform comprises one or more sensor elements. A sensorelement is a component of the sensor platform that is capable ofrecognizing or reacting with an analyte whose presence is to be detectedby the sensor platform. Typically, the sensor element produces adetectable signal after interacting with the analyte to be sensed via anelectrode in the sensor platform, for example. Individual sensorelements within the sensor platform can sense the same or differentanalytes. In this context, the sensor platform can be adapted to measuremultiple analytes simultaneously. For example, multiple individualsensor elements adapted to sense different analytes can be exposed tothe external environment at the same time. Alternatively, multipleindividual sensor elements adapted to sense different analytes can beexposed to the external environment at different times. Otherembodiments include a sensor platform adapted to function asmulti-analyte sensor on a single chip (or, alternatively, on multiplechips). In certain contexts, a signal from an individual analyte sensorelement within the plurality of analyte sensor elements that contact andsense an analyte in a sensor platform are individually interrogatedand/or read. Alternatively, multiple analyte sensor elements within aplurality of analyte sensor elements that contact and sense an analytein the sensor platform are interrogated and/or read simultaneouslyand/or in combination.

In one embodiment, the sensor element utilizes an enzyme (e.g., glucoseoxidase (GOx)) that has been combined with a second protein (e.g.,albumin) in a fixed ratio (e.g., one that is typically optimized forglucose oxidase stabilizing properties) and then applied on the surfaceof an electrode to form a thin enzyme constituent. In one embodiment,the sensor element comprises a GOx and HSA (Human Serum Albumin)mixture. In this embodiment, the GOx reacts with glucose present in thesensing environment (e.g., the body of a mammal) and generates hydrogenperoxide, wherein the hydrogen peroxide so generated is anodicallydetected at a working electrode in the sensor platform.

An “electron transfer agent” is a compound that carries electronsbetween an analyte and a working electrode, either directly, or incooperation with other electron transfer agents. One example of anelectron transfer agent is a redox mediator.

The measurement of analytes including glucose, lactate, etc., isachieved using an external unit and an implantable biosensor platform.The external unit can provide controls for sensor unit selection andoutput display. In one embodiment, the device integrates sensors withbiocompatible coatings as well as drug dispensing devices. In anotherembodiment, the device is capable of additional wireless communicationwith health service providers as appropriate.

The device of the present invention comprises a sensor platform and anexternal unit that are in operable communication through a set oftransceivers. A transceiver comprises an optical transmitter and anoptical receiver. In one embodiment, the optical transmitter is alight-emitting or laser diode. In another embodiment, the opticalreceiver is a photodetector. The two components of a transceiver arelocated in the external control unit (the optical transmitter) and theimplanted sensor platform (the optical receiver), respectively or viseversa. The transceiver orientation is defined by the direction oftransmitted light. The interactive coupling between two transceivers(transmitting in opposing directions) establishes a feedback loop viaother circuits. Two transceivers with opposing location of transmittersand receivers form a closed-loop, capable of wirelessly transmitting andreceived commands, carrying out certain instructions as well astransmitting certain information back and forth among the two units.This interactive feedback loop enables the remote operation of thesensor platform. In addition, with the use of logic and routingcircuits, the feedback loop provides multiple functionalities includinginitialization, calibration, and measurement of one or more analytelevels.

In one embodiment, the external unit comprises an optoelectronicreceiver suitable for receiving optical pulses from the sensor platform,and converting these optical pulses to electrical pulses. In addition,the external unit contains integrated circuits suitable for processingand displaying the analyte levers that are coded in terms of pulsecharacteristics. An optical source located in the external unit powers aplurality of photovoltaic cells, which in turn serve as the power sourcefor the implantable sensor platform.

In one embodiment, the sensor platform comprises a power source; one ormore electrochemical sensor elements suitable for sensing of one or moreanalytes; one or more interfacing circuits providing operating voltagesand a reference voltage to the sensor elements, wherein the interfacingcircuit generates a signal proportional to the amount of analytepresent; (e.g., an ultrasound transmitter) one or more signal processingcircuits in operable communication with the interfacing circuit, whereinthe signal processing circuit converts the analog sensor signal todigital pulses, one or more electrical to optical converters in operablecommunication with the signal processing circuit, wherein the electricalto optical converter converts the digital pulse to optical pulses; and atransmitter for transmitting the optical pulses to the external unit. Inone embodiment, the interfacing circuit comprises a potentiostat. In oneembodiment, the electrical to optical converter is an infrared (IR)transmitter suitable for the wireless relaying of analyte levels andpower management information to the external unit.

In one embodiment, the sensor platform comprises three sub-chips. In oneembodiment, sub-chip #1 comprises a wireless photovoltaic powering solarcell array to power all components of the sensor platform, aphotodetector (PD_(M)) to monitor the power level, an infra redtransmitter (TX_(D)), and a photodetector (PD_(SS)), along with theirassociated circuits. Sub-chip #1 preferably faces the portion of theexternal unit that serves as the power source to power photovoltaiccells (e.g., super luminescent LEDs or laser diodes). The PD_(SS) putssub-chip #1 in operable communication with sub-chip #2. For example, thephotodetector (PD_(SS)) interfaces with a Mode Selector circuit block onsub-chip #2, which in turn communicates with Router/Logic/Mux circuits.Information regarding power levels is ensures that the desired voltageand current levels are available to operate all electronic andoptoelectronic circuits of internal implantable platform unit. This canprevent faulty internal unit readings due to voltage-current levelsbelow threshold. The infrared transmitter TX_(D) also serves to transmitinformation to the external unit regarding the photovoltaic power level.

Sub-chip #1 can further comprise an eye-safe infrared (IR ˜1.55micrometer) InGaAsP—InP LED/laser source (TX_(D)), for example, bondedonto aSiO₂ coated Si substrate in the vicinity of the solar cell array.An 1.55 micrometer IR detector (PD_(D)), located in the external unit,detects the coded internal unit signal. In an alternative embodiment,the external unit further comprises a band-pass filter to rejectradiation from the powering LEDs that operate in a spectral regime,which affords minimum absorption.

Sub-chip #2 comprises one or more interfacing circuits, one or moresignal processing circuits, and one or more electrical to opticalconverters. According to one embodiment, sub-chip #2 comprises a ModeSelector and Router/MUX/Logic blocks, which interface with programmablepotentiostat and calibration circuits, along with a signal processinganalog-to-digital-converter (ADC) interface and TX_(D) driverelectronics. For example, once a sensor is selected, the programmablepotentiostat provides appropriate voltage values for working (V_(W)),reference (V_(REF)), and counter (V_(C)) electrodes of the selectedsensor, located on the Sub-Chip #3. The analog output of the selectedsensor is thus connected (via Router/Logic/MUX block) to thepotentiostat and ADC signal-processing unit. The digital output from theADC circuit is fed to the TX_(D) driver, which in turn is designed tointerface with the infrared transmitter (TX_(D)) on sub-chip #1.

In one embodiment, the analog current developed in a glucose sensor(e.g., due to the presence of glucose in the environment adjacent to theimplanted sensor platform) is converted into voltage pulses of varyingwidth by the ADC circuit. These pulses in turn drive an infrared emitter(TX_(D)). The emitter output is received by an external photodetector(PD_(D)) located in the external unit, which can be worn on the wrist orlocated in a modified PDA unit. Thus, the pulse duration or frequencycarries the glucose level information to the external unit, where it isprocessed and displayed accordingly.

Sub-chip #3 comprises one or more electrochemical sensors, for example,a glucose sensor, along with other micro-sensors (e.g., oxygen, pH,insulin, and ion concentration). In one embodiment, sub-chip #3comprises an electrochemical sensor with working platinum and auxiliaryplatinum electrodes in an inter-digitated configuration, and a referencesilver/silver chloride electrode meandering between the two platinumelectrodes. Sub-chip #3 optionally comprises ionic sensors, in whichfield-effect transistors with an electroactive gate material coating areused.

FIG. 1 illustrates schematically an embodiment of the functional blocksof both the external control unit (1) and the sensor platform (2)subcutaneously implanted under the skin (3).

The external control unit (1) comprises a microprocessor (11), asoftware interface or program (12), a mode select comprising variousswitches (13), and various electronic and optoelectronic “Add-on Devicesand Control Circuits” (14). In addition, there is a display (15) andprovision for interface with “Other Devices” (16). The Add-on devices(14) include an optical source (“719 nm Laser/LED (A)) or sources atwavelengths that are not absorbed by the skin and subcutaneous tissuefor powering solar cells (41) located on Sub-Chip #1 (4) of theimplanted sensor platform (2). The Add-on devices (14) also includes atransmitter (TX_(SS)) (18), operating in the spectral range 800-980 nm,which sends optical commands as coded pulses to the PD_(SS)photodetector (44), located on Sub-chip #1 (4) of the implanted sensorplatform (2). The Add-on devices (14) also includes a photodetector(PD_(D)) (19) operating at 1.55 μm, which receives information as codedoptical pulses from the transmitter (TX_(D)) (45) located on Sub-chip #1(4) of the implanted sensor platform (2). An optical filter isoptionally placed in front of photodetector PD_(D) (19) in order toallow transmission of wavelengths of 1.55 μm and reject away shorterwavelength radiation.

FIG. 1 also shows an embodiment of the implantable sensor platform (2)having a compact size of 0.5 mm width×5 mm length×0.5 mm height. In thisembodiment, the power source of the internal unit comprises photovoltaic(PV) solar cells (41), which are powered by high efficiencylight-emitting diodes (LEDs) (17) in the external unit. These PV cells(41), operating at designed wavelength that allows transmission throughthe skin, provide sufficient power output (voltage and current) neededby the electronic and optoelectronic devices of the implantable sensorplatform unit (2).

In this embodiment, the implantable sensor platform (2) comprises threesub-chips with the following functionality: Sub-chip #1 (4) comprisesthe power source (41), the power level monitor photodetector (PD_(M))(42), the optical command receiver photodetector (PD_(SS)) (44) alongwith its band-pass filter (BPF) (43), and the transmitter (TX_(D)) (45)for transmitting the optical pulses to the external unit (1); Sub-chip#2 (5) comprises the Mode Selector circuit block (51), which interpretsthe optical commands from the transmitter (TX_(SS)) (18) to thephotodetector (PD_(SS)) (44) and communicates it via electrical digitalpulses to the Router/Logic/MUX circuit block (52). The Router/Logic/MUXcircuit block (52) interfaces with the programmable Potentiostat (54),Calibration and Initialization Circuits (53), signal processing circuits(Analog-to-Digital Converter (ADC)) (55), and TX_(D) Driver circuit(56). The The Router/Logic/MUX circuit block (52) along with theprogrammable Potentiostat (54) interfaces with various sensor elementslocated on Sub-chip #3 (6); and Sub-chip #3 (6) comprises an number ofelectrochemical sensors, whose share the same reference (61) and counter(62) electrodes. Three working electrodes (63), (64), and (65) areexplicitly shown on Sub-chip #3 (6). In alternative embodiments, theinternal implantable sensor platform unit (2) comprises two sub-chips oreven one chip, if integration of circuits is miniaturized further.

In operation, the implantable sensor platform unit (2) receives thepowering light (31) from the optical source (17) through the skin (3).This powering light (31) is received by photovoltaic (PV) cells (41)that provide power to all devices and circuits in the implantable sensorplatform (2) through bus lines (23), shown in bold. This powering light(31) is also received by the PD_(M) photodetector (42), which providesinformation regarding the input power level and hence the power producedby the PV cells (41). The implantable sensor platform (2) also receivesthrough the PD_(SS) photodetector (44) optical command and controlinformation (from the external unit via TX_(SS) (18)) as pulses in acertain frequency range (f₁). These optical commands enable selection ofvarious function of the implantable sensor platform such asinitialization, sensor selection, calibration and measurement of analytelevels, power level check, etc. These functions are carried out by theMode Selector (51) and Router/Logic/MUX circuit (52) blocks. These twounits provide interface with all other electronic and optoelectronic andelectrochemical devices and circuits within the implantable sensorplatform unit (2). The PD_(SS) photodetector (44) has a band pass filter(43) or a coating that blocks the incoming powering light (32) (fromoptical source (17)) and reflects it away (33). This preventsundesirable interference of powering light (32) with the optical pulses(34).

Once a function (e.g., initialization, sensor selection, calibration ormeasurement) is selected (through the microprocessor (11) and associatedSoftware (12)), the Mode Select Unit (13) in the external control unit(1) sends encoded electrical pulses, which are transmitted optically bytransmitter TX_(SS) (18) to the implantable sensor platform (2) wherethey are received by the photodetector PD_(SS) (44) and processed by theMode Selector (51) and Router/Logic/MUX (multiplexer) block (52). Uponexecution of a selected function, the result is transmitted to theTX_(D) driver (56), which in turn powers the TX_(D) optical transmitter(45). The TX_(D) transmitter (45) relays the information via opticalpulses (35) of a different frequency range (f₂) through the skin (3) tothe PD_(D) photodetector (19) located in the external unit (1). Thissignal is then processed by the Add-on Devices & Control Circuits (14)of the external control unit (1) in conjunction with the microprocessor(11) and the program loaded in the Software Interface (12) unit. Thesesteps constitute a feedback loop to interactively implement a function.This loop is repeated for every function including initialization,sensor selection, calibration and measurement described below.

An exemplary initialization function operates as follows. Theinitialization function checks if the solar cells are receiving adequateoptical power from the optical source (17). For this, the Mode Selector(51) in conjunction with the Router/Logic/MUX block (52) compares theoutput of the PD_(M) photodetector (42) using a comparator with apredetermined reference, available in the Calibration Circuit unit (53).If the power level is adequate or inadequate, a signal is transmittedusing the TX_(D) driver (56) and TX_(D) transmitter (45) to the externalunit to take the appropriate action (i.e. if power is adequate proceedwith the next function or if the power level is not appropriate,increase the power level of the optical source (17) via the circuits inthe Add-on Devices & Control Circuits unit (14).

An exemplary sensor selection function operates as follows. The sensorselection function, a command comprising an optical pulse set istransmitted by TX_(SS) (18) to the PD_(SS) (44), which selects one ofthe three working electrodes (63, 64, and 65) shown in Sub-chip #3 (6).Once a sensor is selected, “sensor calibration” is typically performed.Sensor calibration includes configuring a programmable potentiostat (54)such that the voltage between the working electrode (63 or 64 or 65)with respect to reference electrode (61) is at the desired valuedictated by the electrochemical reaction involving the detection of acertain analyte. Configuring of the potentiostat determines theappropriate voltage or current levels, as well as the mode of operation(continuous or pulsed for certain duration). This configuration isachieved by Mode Selector (51) and Router/Logic/MUX (52) circuit blocksin conjunction with the Potentiostat (54) and Calibration Circuits (53).Once the optional calibration is accomplished, the next function issensor reading. This function is performed using potentiostat (54) andthe Signal Processor & ADC block (55). The digital output of the SignalProcessor & ADC block (55) is then fed to the TX_(D) driver (56), whichin turn powers the optical transmitter TX_(D) (45). The analogelectrochemical current generated by the potentiostat-driven sensor[which includes three electrodes: a working (63 or 64 or 65), a counter(62) and a reference (61)] is read, amplified, and digitized by theSignal Processor & ADC block (55). The Signal Processor & ADC block (55)converts the magnitude of the electrochemical current into digitalpulses whose frequency is proportional to the analyte level. The digitalelectrical pulses are converted into digital optical pulses that aretransmitted by TX_(D) (45). These optical pulses (35) pass through theskin (3) and are converted back to electrical pulses by photodetectorPD_(D) (19). These electrical pulses are decoded by the external controlunit (1) using the Add-on Devices and Control Circuits (14), and theanalyte level is displayed on Display (15).

Changing from one sensor to another is accomplished, for example, byre-programming of the Router/Logic/MUX (52), which in turn reconfiguresappropriately the Programmable Potentiostat (54). The Router/Logic/MUX(52) selects one of the desired working electrodes (63, 64, or 65). Allof these commands are executed from instructions transmitted using thetransmitter TX_(ss) (18) and its complementary photodetector (PD_(ss))(44) in the implantable sensor platform (2). Their signals are processedby the Mode Selector (51), which is interfaced to the Router/Logic/MUX(52). Router/Logic/MUX (52) performs the reconfiguration and connectionto the calibration circuits. The results of calibration and comparisonare fed through Router/Logic/MUX (52) to the TX_(D) Driver (56) andTransmitter TX_(D) (45) and relayed to the external control unit tocomplete the instructional set and desired function.

FIG. 2 illustrates an exemplary external control unit (1) andimplantable sensor platform (2) comprising three subchips (4, 5, and 6)that are coated with a biocompatible coating (68) containing, forexample, a number of tissue response modifying agents. In oneembodiment, sensor platform (2) is implanted subcutaneously under theskin (3) with a hypodermic needle (69) outfitted with a plunger (70).Implantation takes place, for example, by lifting the skin up, insertingthe needle containing the sensor platform (2) along with itsbiocompatible coating (68) and its plunger (70). After positioning thehypodermic needle (69) to the proper depth, the plunger (70) is heldfixed and the needle (69) is removed, leaving the sensor platform withits biocompatible coating at the desired location. Finally the plungeris also removed. Care is exercised to ensure that the photovoltaic cells(41) are facing up towards the skin (3). FIG. 2 also schematically showsthe subchip construction, interconnections and optical powering andcommunication interface with the external unit (1). Here, only theoptical modules [laser/LED (17), TX_(SS) (18) and PD_(D) (19)] of theexternal control unit (1) are shown. Some interconnects are labeled(21). Working electrodes are shown with specialized coatings (66 and67), specific to the specific analyte sensing.

FIG. 3 illustrates a three-dimensional (3D) schematic representation ofthe 0.5×5 mm implantable wireless platform, consisting of threesub-chips that are encapsulated within a biocompatible coating (34),containing a variety of tissue response modifiers (TRMs) to control, forexample, tissue response and induce neo-angiogenesis. Sub-chip #3 (6)consists of an electrochemical sensor interfaced with sub-chip #2 (5)comprising a potentiostat (54), ADC (55) and IR driver (56) (for glucosetransmitter TX-D (56)). Sub-chip #1 (4) includes the powering solarcells and IR transmitters. The external command, control and monitoring,modified-PDA unit is shown above the implantable sensor capsule. Thisunit consists of powering-LEDs (17) and photodetectors (19) for glucoselevel display (15) and power control circuits (14 in FIG. 1),respectively. In one design embodiment, the implantable device comprisesthree sub-chips, (each of about. 0.5 mm wide, 5 mm long), which arestacked on top of each other to achieve a compact configuration,suitable for needle-assisted administration (see FIG. 2). Sub-chip #1(top) comprises the solar cells (41 in FIG. 2), photodetectors andtransmitters (44) (including PD 42 shown in FIG. 2). Sub-chip #3 (6,bottom chip) comprises the three-electrode (working, reference andcounter) electrochemical glucose sensor. Sub-chip #2 (5, middle)comprises a potentiostat (54), which is operably linked by vias (301between subchip #1 and subchip #2, and 302 between subchip #2 andsubchip #3; see also FIG. 10) to the electrochemical sensor (61-65 inFIG. 1) on subchip #3 and the solar cells on subchip #1 (4), generatingan amperometric current that is fed to the signal-processing unit (55).This unit comprises a transconductance amplifier and ananalog-to-digital (ADC) converter to convert amperometric current (i.e.,glucose current) to voltage pulses. These pulses are fed to a driver(56) which in turn feeds an infra-red 1.55 μm transmitter (TX) (45)(located on sub-chip #1 (4)), using a low-power laser/LED. The opticalpulses that carry glucose level information (defined as TX_(D), D fordisplay) are received by a photodetector (PD_(D)) (19/15) located in theexternal unit. [Various components of sub-chips are shown in FIG. 9].

One feature of the device is an interactive feedback system between thesensor platform and the external unit. This feedback system provides theability to, for example; verify adequate power levels; account for othermeasurements such as blood pressure, body temperature, as well asfactors such as pH and oxygen level, which assist to check biointernalunit calibration; and select a sensor and retrieve the information.

For example, one feedback system provides information regarding whetherthe LEDs (in the external unit) are powered in a manner that ensuresadequate optical input to the solar cells of the sensor platform (fordetails see FIG. 13). The solar cells in turn are supplying variousunits of the internal unit adequate voltage, current and power levels.One scheme to ensure this is to have a feedback system consisting of aphotodetector (PD) monitor measuring the level of optical power receivedfrom the LEDs in the external unit(for details see FIG. 13). The PDoutput current is proportional to the light intensity received. Thiscurrent is converted into a voltage, which compared in a comparator(COMP) against a reference. This reference may be derived from the samecircuit, which supplies the reference electrode of the internal unit.The difference between the reference voltage (V_(REF)) and the PDderived voltage is fed to the infrared transmitter (IR-M in FIG. 3 andTX-M in FIG. 1). This signal is transmitted to the external unit. Thereceived PD input level is compared in the CTRL unit, which in turnadjusts the power level of the LEDs.

The device can further comprise another feedback system regardinginformation about other measurements by the internal unit, such as bloodpressure, oxygen and pH levels, which can affect the calibration oraccuracy of the internal unit.

In another embodiment, the device comprises a plurality of sensorsoberably (operable/fluid/electrical/optical) in communication with theexternal unit. The operative connectivity can be provided by anycommunication means, such as fluid, electrical, optical, or acombination of at least one of the foregoing. The plurality of sensorscan be housed within a single internal unit or multiple internal units.Once a user selects a particular sensor using the sensor select switchin the external unit, it activates an IR transmitter in the externalunit. This transmitter communicates and connects to the associatedphotodetector (PD) shown schematically in FIG. 1 along with the blockmarked VDD, Vref and Control. The communication code enables the PD tointerface with the potentiostat to set desired working and referencevoltages needed to operate the selected sensor. In addition, theinformation is provided to the router switch, which physically selectsthe sensor and connects the potentiostat output to its terminals.

FIG. 4 shows a schematic representation of the optical powering andcommunication interface located in the external control unit (1). Thisinterface also contains a miniaturized camera (71), which assists inlocating and aligning the implantable sensor platform (2) with respectto the external control unit (1). The optics elements within theexternal control unit (1) are housed within the casing (76) and includethree cube reflectors (73, 74, and 75) equipped with the appropriatecoatings to stir reflections depending to the wavelength of theradiation, along with a number of focusing lenses (72), transmitters,receivers (photodetectors) and various other beam shaping components.The light (31) from the optical powering source (17) is collimated bylens (72), steered by two cube reflectors (73 and 74) to illuminate abroad area around the sensor platform (2). This light also providesillumination for the miniaturized camera (71). The 719 nm lightpenetrates the skin adequately and provides enough contrast to theoperator to locate and align the control unit (1) with respect to sensorplatform (2). Optical commands from the TX_(SS) transmitter (18) aredirected through a lens (72) and two cube reflectors (75 and 74) toilluminate the similar broad area around the implanted sensor platform(2). The optical signal from the TX_(D) transmitter (45), located in theimplantable sensor platform (2) passes thought the cube reflector (74)and deflected by the mirror of the cube reflectors (75), collimatedthrough lens (72) and detected by the photodetector PD_(D) (19). Theinsertion of long wavelength pass filter (10) ensures that nointerference from the 719 nm (17) and 800-950 nm (18) light sourcesoccurs. Similarly, the band-pass filter (43) on top of the PD_(SS) (44)detector in the implantable sensor platform (2) ensures no interferencefrom the 719 nm (17) and the 1.55 μm (45) sources, respectively.

FIG. 5 illustrates the selection of one sensor (80 or 81) from animplantable sensor platform that comprises more than one sensor (80,81). It also demonstrates the operation of a programmable potentiostatconsisting of operational amplifiers (87) and associated resistornetwork (82) assisting in the determination of the appropriate voltagebetween the Reference (R, shown in 80 or 81) and the Working (W, shownin 80, 81) electrodes of the selected sensor. A set of coded optical(850-980 nm) pulses (34) (see FIG. 2) transmitted by the TX_(SS) (18),located in the modified-PDA unit (101) are received by the photodetectorPD_(SS) (44). The pulses enable selection of a function (such aspotentiostat reference check, sensor selection, calibration or reading,or checking solar input power level etc.). These are enabled with thehelp of sensor select block 51, switches (shown as 82, 83, 84; andrealized by transistors in integrated circuit chip), circuits lumped inblock 85 and Mux 86. The outcome of this function is communicated backto the modified-PDA using the 1.55 μm optical pulse transmitter (45,located at bottom right corner). Using logic/router block 85 (consistingof reference comparator and calibration circuits), we can provide theinformation to MUX (Multiplexer) circuit (86), which in turn connects,to the driver (56) of TX_(D). The use of different wavelengths and pulsefrequencies for the two transceivers (one for sensor select and theother for rest of the functions) will further ensure minimalinterference. Here, the sensors (80,81) interface with the currentmirror 88, and signal processor 89, and Mux 86, and the driver 56, andthe transmitter TX_(D) (45). The current mirror (88) processes thesensor current and the ADC (89) converts into electrical pulses ofdifferent frequencies depending on the sensor-produced current level.The driver (56) makes it suitable to drive the output light emittingdevice (45, TX_(D)). In the case of ADC, we show an operationalamplifier (501), a comparator (502), One shot (monostable multivibrator,503), and a Divider by 2 circuit (Div 2, 504). There are many variationsto implement this.

In one embodiment, two sensors (S1 and S2) are connected to aprogrammable potentiostat whose reference voltages will be selected onthe sensor under test. The signal processing units can be miniaturizedby reducing the design rules (fineness of microelectronic features) from0.35 μm to 0.12 μm and below.

In one embodiment, two sensors (S1 and S2) are connected to aprogrammable potentiostat whose reference voltages will be selected onthe sensor under test. The control lines and associated switches (to bereplaced by FETs) in the FIGS. 5 and 6 are used to perform selectfunction (including sensor selection S1 or S2 (see FIG. 7)) or CP1 andCP2. The signal processing units can be miniaturized by reducing thedesign rules (fineness of microelectronic features) from 0.35 μm to 0.12μm and below.

In one embodiment, the device is run in self-calibration mode. Thepotentiostat reference voltage [CP1 or CP2 (Potential Reference Check 1or 2) by comparing the voltage difference between the reference andcorresponding working electrode], power level and/or voltage out put ofthe solar cells that are powering the entire chip and transceivers canbe checked when the device is in self-calibration mode. Built-incomparators and logic are used to achieve the self-calibrationfunctions. In operation, a comparator receives a specific signal andcompares it with the reference (voltage) and depending on the differencegenerates a decision, which is then executed via the logic circuits. Forexample, the circuit block labeled “Ref. Comparator and CalibrationCircuits” (85) along with the multiplexer (Mux) (86) enable utilizationof the 1.55 μm transmitter to report back to the PDA the desiredinformation. (FIG. 5)

FIG. 6 shows an alternate scheme wherein a PDA unit is interfaced viawireless technology (e.g., Bluetooth®) to an auxiliary external unit(91), which communicates with the implanted sensor. The signal (from PDAor modified PDA (101)) is received by unit 91, this signal is convertedinto optical pulses by an optical transmitter TX_(SS) (18). This signalis received by photodetector PD_(SS) (44) in the implanted unit. Theinformation from the implanted sensor is received by photodetector PDM(42) now located in the auxiliary external unit (91) which in turntransmits it back to the PDA (101). This type of arrangement isenvisioned in circumstances where a PDA unit is not a dedicated unit formonitoring analytes.

FIG. 7 shows an embodiment of a design of sensor select circuit. Thecircuit has a photodetector (shown as diode on the extreme left) PD_(SS)(44). This photodetector receives optical pulses (800-950 nm) from thetransmitter TX_(SS) (18) located in the external unit (see FIG. 2). Thesignal is amplified by circuit shown in block (703) which consists of aninverting amplifier (two transistors) and a tri-state buffer (702).Block (704) is the timer that generates clock pulse based on our logicblock (705) operation. Depending on the pulse code, one of the sixoutputs (labeled as 707, 708, and 709) gets activated. The activationdepends on block (706, consisting of D-Flip Flops) and Logic Block(705). For example, if the pulse code is set for selection of Sensor #1left most output SN1, shown as part of (707), is selected. This means ahigh voltage is available, which in turn will activate switches 82, 83,and 84 (see FIG. 5). This selection also activates the Router/Logic/Mux(FIG. 1, FIG. 5) circuit in a certain way to enable the activation ofTX_(D) (45) through driver (56). The transmitter TX_(D), operating at1.55 μm, sends a signal showing the selection of Sensor 1. Details oflogic (705) and MUX (86 in FIG. 5) circuits are not shown here.

When Mode select (13, FIG. 1) sends a code for calibration check ofpotentiostat, CP1 and CP2 (shown as 709) are activated. Similarly, modeselect (13) can select calibration control to check power level of solarcells (41) using CSP-I and CSP-0 (708). In FIG. 5, the solar power checkis written as CSC (in block 51).

Circuit designers may implement this concept in a variety of ways.

FIG. 8 shows photographs of an embodiment of a working glucose sensorand the data obtained therefrom. FIG. 8(A) shows an ADC signal processor(MOSIS fabricated chip) interfaced with a hybrid potentiostat. FIG. 8(B)shows the digital pulses on an oscilloscope along with the potentialdifference (680 mV shown on a meter) between the working and referenceelectrode of a functional glucose sensor. The electrochemically detectedglucose concentration is encoded in terms of digital pulse frequency bythe chips in FIG. 8(A). FIG. 8(C) illustrates a plot of glucoseconcentration as a function of digital pulsed frequency, generated bythe signal processing chips shown in FIG. 8(A).

FIG. 9 shows first and second generation microelectronic sensorcomponent embodiments. The dashed-line square and circle indicate thefirst and second generation components, respectively. Selectedcomponents (1-5) are further magnified in panels (B-F) along with theiractual dimensions. (B) shows a 1^(st) generation (MOSIS—fabricated)integrated circuit that contains an ADC signal processor unit and two RFantennas. (C) shows a planar electrochemical sensor with threeelectrodes (working (Pt), counter (Pt) and reference (Ag/AgCl,whitish-appearing, right-most electrode)). (D) shows an InGaAsPsemiconductor laser stack (8 in number) emitting infra red (IR)radiation of 1.55 μm. The sensor platform may be equipped with one ofthese lasers. (E) shows an individual, Si-based solar cell device.Typically 5 or 6 panels may be connected in series to power up allcomponents of the implantable sensor platform. (F) shows a secondgeneration (MOSIS—fabricated) integrated circuit containing thepotentiostat and ADC signal processors, without the RF antennas.Stacking of all components from the second generation will result inabout a 0.3×5 mm implantable platform with provision for 0.1 mm thickbiocompatible coating on each side. The overall 0.5×5 mm sensor can beimplanted through a 16 gauge hypodermic needle.

FIG. 10 shows an exemplary layout of three sub-chips using a hybridapproach to integrate sub-chips with robust interconnects 1003 betweensubchip #1 and subchip #2 (labeled as 112G for ground interconnectbetween subchip #1 and subchip #2, 112P for power or voltage supply) andinterconnects 1004 between subchip #2 and subchip #3 (e.g. labels I23W,I23R and I23C; here W for working electrode, R for reference electrodeand C for counter electrodes connections between sensors on subchip #3and potentiostat on subchip #2). The subchips are connected using vias(1001) and bumps (1002). The vias are V12P, V12S, V12G (numeral 12 referto subchips #1 and 2, and the letter P, S, and G refer to function suchas power, signal and ground). The bumps 1002 shown on subchip #3 matewith the vias on the subchip above this (that is subchip #2). Bumps aredesignated as B12P (the bump on subchip #2 locks with the via on subchip#1; here P refers to power or voltage connection) and B12G (G for groundconnection). Cross-sectional schematic illustration of bonding between avia and bump for sub-chips #1 and #2 (bottom left). Top view ofinterconnect and power pad are also shown. Bonding between sub-chip #2and sub-chip #3 is shown using two approaches used in the integration(bottom right). In FIG. 2 the word ‘vias’ (301 and 302) was used for thecomplete connection between two subchips including via, interconnect andthe bump.

FIG. 11 shows another embodiment of a three sub-chip design (shown inFIGS. 1, 2 and 3) Here the top surface of Sub-chip #1 (4) ishermetically sealed using a transparent glass window (111) sealed toraised silicon walls (110) via anodic bonding (112) between the glassand the silicon. Similar hermetic seals are employed between the bottomsurface of Sub-Chip #1 (4) and the top surface of Sub-Chip #2 (5), aswell as the bottom surface of Sub-Chip #2 (5) and the back surface ofSub-Chip #3 (6). The front surface of Sub-Chip #3 (6) hosts variouselectrodes that are designed to operate in body fluids. Allinterconnects among sub-chips are vias filled with non-corrosive metalsthat provide hermetic seals for microelectronic and optoelectroniccomponents of the implantable sensor platform.

FIG. 12 shows an embodiment where the three sub-chip design has beenimplemented as two sub-chips. This design is possible using finer designrules (i.e., 0.35 micron versus 0.5 micron chip processing) for thepotentiostat and signal processing components (usually located onsub-chip #2, FIG. 11), which results in saved space on a Si chip. As aresult, functions of sub-chip #2 and sub-chip #3 could be integrated asone sub-chip (referred to as (120) the new-sub-chip #2); thus a sensorplatform having two sub-chips can be realized. An exemplary embodimentis illustrated in FIG. 12. New-Sub-chip #2 (120), according to thisembodiment, comprises a ˜0.25×4 mm MOSIS chip (122) (like part 5 of FIG.1, or electrical circuit of FIG. 5) bonded face down on the carrier0.3×5 mm platform (120) that hosts along its perimeter theelectrochemical sensor electrodes (61,62,63,64). The bonding pads (123)are connected to the bump and or vias (shown as 1002 and 124). Twoworking electrodes (63,64) are placed at either side along the width ina meander-form to improve adhesion of the glucose oxidase coating.Lengthwise, a Pt counter (62) and an Ag/AgCl reference electrode (61)are also placed to function in conjunction with either workingelectrode. In addition, a 1.55 μm InGaAsP transmitter TX_(D) (45) (andTX_(M) which is for monitoring and could be one unit via MUX (18)circuits) is placed on the upper right corner, adjacent to the workingelectrode 2 (64). The wafer carrier is a high resistivity (>20,000Ohm/cm) Si (100) in which bumps, vias and interconnects are patternedand metallized prior to growing the organic layers for theelectrochemical sensor. The bumps and vias are also shown. Both theMOSIS and the 1.55 μm InGaAsP transmitter chip are placed face down andtheir pads are connected to the pads on the carrier chip.

The subchip #1 (4) is shown as part new subchip #1 (121). In thisembodiment, new-sub-chip #1 has 3 pads on either side for power supplydistribution (e.g., V_(dd), V₁ for voltages, and C for common). Thepower is supplied to the new-sub-chip #2 using metalized vias labeled asV_(VDD) (124). In addition, vias are used to connect photodetectorsPD_(SS) (44) and PD_(M) on new-sub-chip #1 to new-sub-chip #2 (havingelectronics such as sensor select, routing logic/MUX, etc.). Note thatthe 1.55 μm transmitter is located on new-sub-chip #2 as this wavelengthis transparent to Si platform and chips. Solar cells (41) are not shownindividually as in FIG. 2.

FIG. 13 shows embodiments of various circuits that may be integratedinto two sub-chips: (subchip #1 (121 of FIG. 12 or 4 of FIG. 1) andsubchip #2 (120) of FIG. 12. It also shows some circuits (14 calledAdd-on Devices & Control Circuits in FIG. 1) that may be housed in themodified PDA unit (101). Here, the modified PDA is communicatingdirectly with the implanted unit.

In some embodiments, it is advantageous to employ an auxiliary externalunit (91) [of FIG. 6] that communicates with a PDA (101) on one hand andthe implanted sensor platform on the other. This is shown in FIG. 14. Inthis embodiment, the PDA unit is communicating with an auxiliaryexternal unit (91) via wireless technology (e.g., Bluetooth®). Theexternal unit communicates with the implanted unit using opticalcommunication.

FIG. 15 shows a methodology to integrate two sub-chips into one waferplatform using wafer bonding methodology. This advanced integration doesnot require vias and bumps as discussed above. In this embodiment, thephotodetectors [PD_(M) (44) and PD_(SS) (42)] and solar cells (41) arerealized on top part of the Si chip (˜20-40 microns in thickness) havingelectrical resistivity in the 1-10 Ohm-cm. This wafer is bonded [usingsilicon dioxide layer (151) or other wafer fusion techniques known inthe literature] to a high resistivity (˜˜10,000 Ohm-cm) wafer. Thiswafer or the bottom side has sensors [comprising of various electrodes(61), (62), (63), and (64)] processed along with signal processing chip(5), which is mounted in a recessed part. The recessed part is createdby deep RIE (reactive ion etching) or other technique. The top part ofthe wafer can be hermetically sealed using anodic bonding using glass111 and bonded seals 112.

FIG. 16 shows a methodology where the internal unit implantable platform(160) is integrated with drug delivery devices (shown as three modulesbelow the command Control Display Wrist Unit (101D)). These units arealso implanted separately or integrated along with the internal unit.Based on the implanted sensor reading of glucose or insulin, the controlunit 101D (e.g. modified PDA unit) sends the signal to the insulindispenser to dispense desired amount of insulin. Here, we have threesub-units or sub-modules identified as subchip #1 (powering modulesimilar to part 4 of FIG. 1), Dispenser subchip #2 (161), and Dispensersubchip #3 (162). Subchip #1 provides the power to unit (161) and unit(162). Dispenser subchip (161) consists of sensor, signal processing [ascarried out by subchip #2 (5) and subchip #3 (6) in FIG. 1], and controlcircuits (163) to actuate MEM (microelectrochemical) actuators (164,165). MEM actuators activate Valves (166, 167) located on the Dispensersubchip #3 (162) which connect to the insulin drug reservoir (168). Theinsulin is dispensed via 173 and 174 if needed at two locations. Theimplanted biosensor in turn provides information regarding the glucoselevels (169) or insulin levels. This information is fed to the processor(170) of the command unit (101D) where it is compared with expectedinsulin levels per software protocol and executed by Level Comparatorunit (171). This in turn activates Drug Delivery Module (172) whichrelays the information to control electronic circuits (163). The commandunit (101D) has optical power source (17) for the solar cells (41) onsubchip #1 (4).

FIG. 17 illustrates glucose sensor with an Ag reference electrode anddetails of coatings on Pt working electrode. This is a schematicrepresentation of modified Clark amperometric glucose sensor, along withvarious chemical, electrochemical and diffusion processes associatedwith its operation. The working Pt electrode (63 or 64 if more than one)is coated with electropolymerized o-phenylddiamine. PPD layer (171)(e.g., prevents permeation of ascorbic acid, acetaminophen etc.) iscoated with glucose oxidase (GO_(x)) layer (172), which in turn iscoated with a semipermeable humic acid membrane (173) to reduce theamount of glucose entering the sensor. An HRP-modified hydrogel layer(174) is then applied to eliminate the outer diffusion of H₂O₂. This isfollowed by an outer composite hydrogel coating with embeddedmicrospheres at different stages of degradation and TRM release (175).The tissue is represented by (176).

In one embodiment, an enzyme-based glucose sensor operates on theprinciple of detection of hydrogen peroxide (H₂O₂) formed by glucoseoxidation. Glucose oxidase (GO_(x)) acts as a catalyst, which turnsglucose into gluconic acid (Reaction 1) and produces H₂O₂. H₂O₂ iselectrochemically oxidized (Reaction 2) under an applied potential of0.7 V and the current measured is related to the glucose concentration(see diagram in FIG. 17). The semipermeable membrane, depicted in FIG.17, in addition to assisting in prevention of biofouling, is designed toregulate glucose diffusion. It is well known that for an enzyme-basedimplantable sensor to work at its optimum efficiency in tissue, theratio of oxygen, which regenerates the enzyme, to the permeating glucoseshould remain constant. Typically, the physiological levels for glucoseand oxygen in subcutaneous tissue are 5.6 and 0.1 mM, respectively.Thus, in the absence of a diffusion-limiting barrier for glucose, thekinetics of H₂O₂ production may be oxygen-limited due to thesignificantly larger amount of glucose compared to oxygen. At highglucose concentrations, this oxygen limit can lead to reduced glucosesensitivity. Therefore, a method to achieve accurate monitoring ofglucose over the entire physiological concentration range with highsensitivity and short response time an outer membrane with tunablepermeability properties is needed. Such a membrane has been developedthrough layer-by-layer grown polyelectrolytes and/or multi-valentcations.

The GOx and/or carrier protein concentration may vary. For example, theGOx concentration is about 50 mg/ml (approximately 10,000 U/ml) to about700 mg/ml (about 150,000 U/ml). In particular, the GOx concentration isabout 115 mg/ml (approximately 22,000 U/ml). In such embodiments, theHSA concentration is about 0.5%-30% (w/v), depending on the GOxconcentration. In particular, the HSA concentration is about 1-10% w/v,and most particularly is about 5% w/v. In alternative embodiments,collagen or BSA (Bovine Serum Albumin) or other structural proteins usedin these contexts can be used instead of or in addition to HSA. AlthoughGOx is discussed as an enzyme in the sensor element, other proteinsand/or enzymes may also be used or may be used in place of GOx,including, but not limited to glucose dehydrogenase or hexokinase,hexose oxidase, lactate oxidase, and the like. Other proteins and/orenzymes may also be used, as will be evident to those skilled in theart. Moreover, although HSA is employed in the example embodiment, otherstructural proteins, such as BSA, collagens or the like, can be usedinstead of or in addition to HSA.

For embodiments employing enzymes other than GOx, concentrations otherthan those discussed herein may be utilized. The concentration may bevaried not only depending on the particular enzyme being employed, butalso depending on the desired properties of the resulting proteinmatrix. For example, a certain concentration may be utilized if theprotein matrix is to be used in a diagnostic capacity while a differentconcentration may be utilized if certain structural properties aredesired. Those skilled in the art will understand that the concentrationutilized may be varied through routine experimentation to determinewhich concentration (and of which enzyme or protein) may yield thedesired result.

FIG. 18 shows the self-assembly of semipermeable membrane composed ofhumic acids and Fe³⁺ on the outer surface of the electrochemical sensor.As the number of bilayers increases, there is an increase in thetortuosity for glucose diffusion towards the enzyme. The glucose oxidase(GO_(x)) layer is coated with a semipermeable humic acid membrane toreduce the amount of glucose entering the sensor.

For testing purposes, a miniaturized sensor (shown in FIG. 9) made ofplatinum evaporated on a high resistivity Si wafer has been developed.To ensure Pt bonding to the Si wafer, a number of coatings have beenemployed to eliminate delamination problems once implantation takesplace. In particular, Au/Ti/Pt/Ti/Ag coatings may be employed. Silver isremoved from the working and counter electrodes. Ag is converted to AgClto form a reference electrode. In the case of the working electrode, afilm of poly (o-phenylenediamine) is electropolymerized on the workingelectrode, following which the sensing enzyme, i.e., glucose oxidase, isimmobilized on top. Finally a semipermeable membrane composed of humicacids and Fe³⁺ ions is grown on the device through electrostaticself-assembly. The sensor is tested via amperometry in phosphatebuffered saline (PBS) solution maintained at 37° C. Glucose is added tothe solution following which a pulse of 0.7V is applied to the deviceevery 5-10 minutes until a constant current reading is obtained. Oncethe device has stabilized, glucose concentration is incremented and theprocess repeated to generate a calibration curve as shown in FIG. 16.

FIG. 19 illustrates pulsed mode operation of the sensor. In pulsed modeoperation, voltages are applied to various sensor electrodes for acertain duration. Calibration curve of current response vs. change inglucose concentration is shown for different glucose concentrations. Asthe amount of glucose in the system is increased, there is acorresponding rise in the current. The Figure on left indicates the timerequired for the device to stabilize at each glucose concentration. Asthe concentration of glucose increases, the device reaches a stablereading faster.

FIG. 20 shows that by varying the number of dip-cycles (2, 5 and 10),the current is reduced by almost 10 fold, while maintaining currentlinearity. The thickness of the film depends on the number ofdip-cycles. The ability of these membranes to act as an efficientbarrier for glucose permeation has thus been demonstrated.

FIG. 21 shows histological evaluations of subcutaneous tissue samplestaken from the vicinity of hydrogel composites containing PLGAmicrospheres at 3 and 21 days post implantation. The representativesections shown are 3 days after implantation (A & B) and 21 days afterimplantation (C & D). There seems to be no inflammation showing theeffectiveness of the coatings.

FIG. 22 shows a schematic cross-section of an embodiment of implantablesensor platform with a methodology to hermetically seal sub-chips usingone high resistivity Si wafer as the carrier (220) which has onerecessed region (221) on one side. In this region solar cells (41) arebonded on metallic pads (224) deposited on an oxide layer (223).Photodetectors (44) and (42) are also placed on this side. The carrierwafer (220) body is recessed (222) on the other side as well. Thinnedsub-chips 2 (5) and subchip #3 (6) are placed in the recessed region(222).

The sub-chips can be electrically interconnected using interconnectslike (226) which run in vias like (225) as shown or other standardinterconnect techniques may be used. The hermetical seal provided by theglass-Si anodic bonding (112) is shown on top as well as bottom surfacesof the carrier (220). This is an alternate approach for three sub-chipintegration as shown in FIGS. 11 and 15.

A variety of optional items may be included in the sensor platform. Oneoptional item is a temperature probe. One exemplary temperature probecomprises two probe leads connected to each other through atemperature-dependent element that is formed using a material with atemperature-dependent characteristic. An example of a suitabletemperature-dependent characteristic is the resistance of thetemperature-dependent element. The two probe leads comprise, forexample, a metal, an alloy, a semimetal, such as graphite, a degenerateor highly doped semiconductor, or a small-band gap semiconductor.Examples of suitable materials include gold, silver, ruthenium oxide,titanium nitride, titanium dioxide, indium doped tin oxide, tin dopedindium oxide, or graphite. The temperature-dependent element can furthercomprise a fine trace (e.g., a conductive trace that has a smallercross-section than that of the probe leads) of the same conductivematerial as the probe leads, or another material such as a carbon ink, acarbon fiber, or platinum, which has a temperature-dependentcharacteristic, such as resistance, that provides atemperature-dependent signal when a voltage source is attached to thetwo probe leads of the temperature probe. The temperature-dependentcharacteristic of the temperature-dependent element can either increaseor decrease with temperature.

The sensor platform comprises components manufactured from biocompatiblematerials, such as materials that are corrosion resistant, INCLUDING Pt,SiO₂ coatings, and glass thin films. In addition, corrosion resistantmaterials that are harmless to tissues in biologic environments, such assilicon and heavily boron-doped silicon can be used in the manufactureof the components of the internal unit. Another method by which thecorrosion resistance of the internal unit can be improved is throughcoating of the internal unit with titanium, iridium, Parylene (abiocompatible polymer), or various other common and/or proprietary thickand thin films.

The sensor platform optionally comprises a biocompatible coating. Thebioactive polymers are generally biocompatible, that is, physiologicallytolerated, and do not cause substantial adverse local or systemicresponses. While synthetic polymers such as poly(tetrafluoroethylene),silicones, poly(acrylate), poly(methacrylate), hydrogels, andderivatives thereof are most commonly used, natural polymers such asproteins and carbohydrates are also suitable. The bioactive polymerlayer functions to protect the implant, preserve its function, minimizeprotein adsorption onto the implant, and serve as a site for thedelivery of the tissue response modifying agents and drugs as well asother drugs and factors.

In one embodiment, the bioactive polymer layer comprises a hydrogel.Hydrogels are formed from the polymerization of hydrophilic andhydrophobic monomers to form gels and are described, for example, inU.S. Pat. No. 4,983,181 and No. 4,994,081, which are incorporated byreference herein. Hydrogels consist largely of water, and may becrosslinked by either chemical or physical methods. Chemicalcrosslinking is exemplified by the free-radical induced crosslinking ofdienes such as ethylene glycol dimethacrylate (EGDMA), and the like.Physical crosslinks are formed by copolymerizing a hydrophobicco-monomer with the water-soluble monomer, and then by contacting thecopolymerized gel with water. Physical association of the hydrophobicregions of the gel results in the formation of physical crosslinks.Control of the ratio of hydrophilic to hydrophobic monomers allowscontrol of the final properties of the gel. Physical crosslinks can alsobe formed by freeze/thaw methods, for example freeze/thawing apoly(vinyl alcohol) (PVA) hydrogel. Highly water-swollen hydrogels arebioactive, and have minimal impact on the diffusion rates of smallmolecules. Hydrogels are also intrinsically mobile, and therefore haveminimal deleterious effects on associated peptide tissue responsemodifiers.

Hydrogels may be formed by the polymerization of monomers such as2-hydroxyethyl methacrylate, 2-hydroxyethyl methacrylate, fluorinatedacrylates, acrylic acid, and methacrylic acid, and combinations thereof.Suitable hydrogels include copolymers of 2-hydroxyethyl methacrylate,wherein the co-monomers are selected to improve mechanical strength,stability to hydrolysis, or other mechanical or chemicalcharacteristics. Copolymerization with various acidic monomers candecrease the buffer capacity of the gel, and thus modulate the releaseof the tissue response modifier. Suitable co-monomers include, but arenot limited to, 3-hydroxypropyl methacrylate, N-vinyl pyrrolidinone,2-hydroxyethyl acrylate, glycerol methacrylate, n-isopropyl acrylamide,N,N-dimethylacrylamide, glycidyl methacrylate, and combinations thereof.Suitable hydrogels are terpolymers of 2-hydroxyethyl methacrylate(HEMA), N-vinyl pyrrolidinone (NVP), and2-N-ethylperflourooctanesulfanamido ethyl acrylate (FOSA) with addedEGDMA to provide controlled crosslinking. HEMA is hydrophilic, andswells in the presence of water. The hydroxyl groups of HEMA alsoprovide potential sites for the covalent attachment of tissue responsemodifiers, slow release delivery systems, and the like. Acrylic acid,methacrylic acid, and other functionalized vinyl monomers can also beemployed to provide these attachment sites. NVP is amphiphilic, whereinthe backbone ring provides hydrophobicity and the polar amide groupprovides hydrophilicity. Poly(vinyl pyrrolidinone) is water soluble,physiologically inactive, and forms complexes with a number of smallmolecules such as iodine and chlorhexidine. Use of NVP improves thetoughness of polymerized HEMA, and provides for the enhanced solubilityof the other monomers under bulk polymerization conditions.

An example of a bioactive layer generated by self-assembly is theformation of NAFION™/Fe³⁺ multilayer films. NAFION™ is a perfluorinatedelectrolyte having sulfonic acid functionalities that has beenpreviously used as a semipermeable membrane for electrochemical sensors.However, the strong ion-exchange properties of NAFION™ lead tocalcification in vitro and in vivo. The sulfonate (R—SO₃) groups presentin the hydrophilic domains of the membrane act as nucleating sites fordeposition of calcium phosphate. These crystals tend to inhibitmetabolite transport through the membrane, and also cause the membraneto become brittle and eventually crack. Electrostatic assembly ofNAFION™ and Fe³⁺ from dilute solutions of ferric citrate at a pH about 2to 6 can be used to prevent calcium deposition.

A natural bioactive coating is a mussel adhesive protein (MAP).Self-assembly of biological materials such as mussel adhesive proteinsallows the incorporation of materials, which improve implantbiocompatibility. MAP produced by the blue seal mussel (Mytilus edulis)generally comprises 75 to 85 repeating decameric units having theprimary sequence of KPSY-Hyp-Hyp-T-DOPA, wherein Hyp is hydroxyprolineand DOPA is 3,4-dihydroxyphenylalanine. DOPA is a strong metal chelatingagent, particularly with Ca²⁺ and Fe³⁺, and the strong self-aggregationof DOPA in the presence of cations results in supra-molecularself-assembly. Accordingly, a substrate comprising metal chelatinggroups, for example free amine groups, is sequentially immersed first ina solution comprising metal ions (i.e., Ca²⁺ and/or Fe³⁺) (followed byoptional washing in fresh solvent); and second, in a solution comprisingthe poly(ligand) (i.e., the MAP protein) (followed by optional washingin fresh solvent). The thickness of the membrane will be directlyproportional to the number of sequential immersion cycles. The assemblyof the membrane may be monitored with Variable Angle SpectroscopicEllipsometry (VASE), UV-VIS and Quartz Crystal Microbalance. The strongchelation between Ca²⁺ and DOPA in the MAP membrane results in asubstantial decrease in porosity, allowing the permeation of smallmolecules such as glucose and oxygen, while excluding permeation oflarger molecules. Additionally, the introduction of small amount ofcrosslinking, via the Michael addition from neighboring lysine repeatsby slight increase of pH above 8.5, which may be used to furtherfine-tune the permeability of such assemblies to levels.

Humic acids may also be polymerized, or self-assembled into abiocompatible layer. Humic acids or “humic substances” areheterogeneous, high-molecular weight organic acids having a largeproportion of DOPA, and are resistant to microbial degradation. Theknown ability of humic acids to donate and accept electrons from avariety of metals and organic molecules explains their capability toshuttle electrons between the humic-reducing microorganisms and theFe(III)-Fe(II) oxide. It has been suggested that humic acids participatein a biological electron transfer as a result of the electron acceptingability of quinone moieties when reduced to hydroquinones andvice-versa. This renders the Fe³⁺/humic acid assembled membranes anattractive vehicle for the attachment to various kind of biocompatiblelayer.

Other components may also be incorporated into the bioactive polymerlayer, such as poly(ethylene oxide) (PEG), to minimize proteinadsorption. Poly(ethylene oxide) is most readily incorporated into thehydrogel, for example, by co-polymerization of a vinyl monomer havingpoly(ethylene oxide) side chains, for example poly(ethylene glycol)methacrylate (which is commercially available from Aldrich ChemicalCo.), or a divinyl-terminated poly(ethylene glycol) macromonomer.Copolymerization of HEMA and poly(ethylene glycol) methacrylate in thepresence of AIBN yields a more flexible, unhydrated copolymer. Theoptimal molecular weight and content of poly(ethylene oxide) for eachapplication can be determined by protein adsorption studies.

To provide further chemical functionality on the bioactive polymerlayer, particularly a hydrogel layer, either polyvinyl alcohol orpolyethylene imine may be employed as macromolecular surfactants. Wherehydroxyl functionalities are available, the coupling is promoted bytresylation. Poly(ethylene oxide) may also be grafted to hydroxyl groupson the surface of the polymer layer by tresylation coupling withJeffamine, an amine-terminated poly(ethylene oxide) commerciallyavailable from Huntsman.

In one embodiment, the biocompatible layer comprises a biocompatiblemembrane, which is permeable to analytes, such as oxygen and glucose,but is impermeable to, for example, white blood cells and macrophages toprevent these cells from contacting other components of the internalunit. The biocompatible membrane can comprise polymers including, butnot limited to, polypropylene, polysulphone, polytetrafluoroethylene(PTFE), and poly(ethylene terephthalate) (PET). The biocompatible layershould be biostable for long periods of time (e.g., several years).

The internal unit can also comprise a mass transport-limiting layer toact as a diffusion-limiting barrier to reduce the rate of mass transportof the analyte, for example, glucose or lactate, into the internal unit.By limiting the diffusion of the analyte, the steady state concentrationof the analyte in the proximity of the working electrode (which isproportional to the concentration of the analyte in the body or samplefluid) can be reduced. This extends the upper range of analyteconcentrations that can still be accurately measured and can also expandthe range in which the current increases approximately linearly with thelevel of the analyte.

In some embodiments, the mass transport limiting layer can also limitthe flow of oxygen into the internal unit. This can improve thestability of internal units that are used in situations where variationin the partial pressure of oxygen causes non-linearity in internal unitresponse. In these embodiments, the mass transport limiting layerrestricts oxygen transport by at least 40%, specifically at least 60%,and more specifically at least 80%, than the membrane restrictstransport of the analyte. In these embodiments, the mass transportlimiting layer comprises a film that is less permeable to oxygen, forexample, by having density closer to that of the crystalline polymer,such as polyesters including polyethylene terephthalate.

FIG. 16 shows a methodology where the internal unit platform isintegrated with drug delivery devices also implanted or integrated alongwith the internal unit.

In one embodiment, the drug delivery device delivers a tissue responsemodifier. “Tissue response modifiers” as used herein are factors thatcontrol the response of tissue adjacent to the site of implantation. Onefacet of this response can be broadly divided into a two-step process,inflammation and wound healing. An uncontrolled inflammatory response(acute or chronic) results in extensive tissue destruction andultimately tissue fibrosis. Wound healing includes regeneration of theinjured tissue, repair (fibrosis), and in-growth of new blood vessels(neovascularization and angiogenesis). For fibrosis, the body utilizescollagen from activated fibroblasts to “patch and fill” theunregenerated areas resulting from trauma and inflammation.

Fibrosis can lead to “encapsulation” or “entombment” of the sensor infibrotic tissue and this can lead to loss of analyte supply and loss offunctionality of the sensor. In-growth of new blood vessels is criticalto the ultimate outcome of wound healing. A number of other responsesare also included within this category, for example fibroblast formationand function, leukocyte activation, leukocyte adherence, lymphocyteactivation, lymphocyte adherence, macrophage activation, macrophageadherence, thrombosis, cell migration, cell proliferation includinguncontrolled growth, neoplasia, and cell injury and death. Adversetissue responses to implantation may also arise through geneticdisorders, immune diseases, infectious disease, environmental exposureto toxins, nutritional diseases, and diseases of infancy and childhood.

Tissue response modifiers are therefore a broad category of organic andinorganic, synthetic and natural materials, and derivatives thereofwhich affect the above responses to tissue injury upon implantation.Such materials include but are not limited to synthetic organiccompounds (drugs), peptides, polypeptides, proteins, lipids, sugars,carbohydrates, certain RNA and DNA molecules, and fatty acids, as wellmetabolites and derivatives of each. Tissue response modifiers may alsotake the form of, or be available from genetic material, viruses,prokaryotic or eukaryotic cells. The tissue response modifiers can be invarious forms, such as unchanged molecules, components of molecularcomplexes, or pharmacologically acceptable salts or simple derivativessuch as esters, ethers, and amides. Tissue response modifiers may bederived from viral, microbial, fungal, plant, insect, fish, and othervertebrate sources.

Exemplary tissue response modifiers include, but are not limited to,anti-inflammatory agents such as steroidal drugs, for examplecorticosteroids such as Dexamethasone(9-alpha-fluoro-16-alpha-methylprednisolone), a potent, broad spectrumsteroidal anti-inflammatory and anti-fibrotic drug with known efficacyin a diabetic rat model, methyl prednisone, triamcoline(fluoroxyprednilisone), hydrocortisone (17-hydroxycorticosterone); andnon-steroidal drugs, for example Ketoprofin (2-(3-benzophenyl)propionicacid), cyclosporin, Naproxin ((+)-6-methoxy-alpha-methyl-2-naphthaleneacetic acid), and Ibuprofin (4-isobutyl-alpha-methylphenyl acetic acid).

Other exemplary tissue response modifiers include neovascularizationagents such as cytokines. Cytokines are growth factors such astransforming growth factor alpha (TGFA), epidermal growth factor (EGF),vascular endothelial growth factor (VEGF), and anti-transforming growthfactor beta (TGFB). TGFA suppresses collagen synthesis and stimulatesangiogenesis. It has been shown that epidermal growth factor tethered toa solid substrate retains significant mobility and an activeconformation. VEGF stimulates angiogenesis, and is advantageous becauseit selectively promotes proliferation of endothelial cells and notfibroblasts or collagen synthesis, in contrast to other angiogenicfactors. In addition to promoting would healing, the improved blood flowresulting from the presence of neovascularization agents should alsoimprove the accuracy of sensor measurements.

Another type of tissue response modifier is a neutralizing antibodyincluding, for example, anti-transforming growth factor beta antibody(anti-TGFB); anti-TGFB receptor antibody; and anti-fibroblast antibody(anti-CD44). Anti-TGFB antibody has been shown to inhibit fibroblastproliferation, and hence inhibit fibrosis. Because of the importance ofTGFB in fibrosis, anti-TGFB receptor antibodies inhibit fibrosis byblocking TGFB activation of fibroblasts. Recent studies havedemonstrated that anti-CD44 antibody induces programmed cell death(apoptosis) in fibroblasts in vitro. Thus, use of anti-CD44 antibodyrepresents a novel approach to inhibition of fibroblast formation, andtherefore fibrosis. Other anti-proliferative agents include MitomicyinC, which inhibits fibroblast proliferation under certain circumstances,such as after vascularization has occurred.

Adhesive ligands (“binding motifs”) may also be used as tissue responsemodifiers, wherein the adhesive ligands are incorporated into thepolymer layer to stimulate direct attachment of endothelial cells toimplant surfaces. Such attachment promotes neovascularization at theimplant/tissue interface. Where the surface density of binding motifshas an effect on the cellular response, variation in the density of thebinding motifs allows control of the response. Exemplary adhesiveligands include but are not limited to the arginine-glycine-asparticacid (RGD) motif, and arginine-glutamic acid-aspartic acid-valine (REDV)motif, a fibronectin polypeptide. The REDV ligand has been shown toselectively bind to human endothelial cells, but not to bind to smoothmuscle cells, fibroblasts or blood platelets when used in an appropriateamount. Sensors detecting body temperature, blood gases, ionicconcentrations can be incorporated in the implantable sensor platform.The analyte sensing device of Claim 1, wherein the sensor elementcomprises a body temperature sensor, a blood pressure sensor, a pHsensor, an oxygen sensor, a glucose sensor, a lactate sensor, or acombination comprising one or more of the foregoing sensors.

In operation, the device can use any mechanism (e.g., enzymatic ornon-enzymatic) by which a particular analyte can be quantitated.

The devices and methods disclosed herein can be applied to determine themetabolic levels of many analytes present in biological fluids,including, but not limited to, glucose, amino acids, and lactate.Suitable analytes include analytes that are substrates for oxidaseenzymes.

Some analytes, such as oxygen, can be directly electrooxidized orelectroreduced on the working electrode. For other analytes, such asglucose and lactate, an electron transfer agent and/or a catalyst canfacilitate the electrooxidation or electroreduction of the analyte.Catalysts can also be used for those analytes, such as oxygen, that canbe directly electrooxidized or electroreduced on the working electrode.For example, some embodiments can quantitate metabolic glucose levels byusing a membrane comprising glucose oxidase (see FIG. 17) that catalyzesthe conversion of glucose and molecular oxygen to gluconate and hydrogenperoxide: Glucose+O₂→Gluconate+H₂O₂. Because for each glucose moleculeconverted to gluconate, there is a proportional change in theco-reactant O₂ and the product H₂O₂, one can monitor the current changein either the co-reactant or the product to determine glucoseconcentration.

In one embodiment, the sensor element comprises an electrochemical pHsensor. Since a large number of biological processes are pH-dependent,there is a great need for outfitting miniaturized biosensors with a pHsensing element. The need for maintaining biocompatibility limits theuse of traditional materials for the fabrication of pH-sensors (i.e.,electrically semiconducting oxides such as MoO₂,⁴³ IrO₂,⁴⁴ or RuO₂ ⁴⁵)due to their toxicity. Biocompatible polymers that contain nitrogen oroxygen moieties amenable to protonation have been used to developpotentiometric pH biosensors. Polyphenol, polyaniline,poly(1,2-diaminobenzene), poly(4,4′-diaminoddiphenyl ether), etc. havebeen employed in fabricating pH sensors. The electrochemistry of thesepolymers however, is greatly affected by redox reagents (such as H₂O₂)and based upon prior experience with poly(o-phenyl diamine), positioningsuch a sensing element in the vicinity of a glucose sensor (whichproduces H₂O₂) could affect the measurements. More recently, linearpolyethylenimine (L-PEI) and linear polypropylenimine (L-PPI) modifiedPt electrodes have been successfully used for the development ofminiaturized electrochemical pH sensors with a linear pH range from 4-9.The non-semiconducting nature of both L-PEI and L-PPI polymers renderthem ideal for operation within a redox-prone environment and theirbiocompatibility and long-term stability (when operated in athree-electrode configuration) renders them ideal for the development ofour miniaturized pH sensors.

In one embodiment, ethylenediamine (EDA) or 1,3-diaminopropane (1,3-DAP)is be electropolymerized onto flat substrate by means of cyclicvoltametry in solutions composed of 10⁻² MN-lithiotrifluromethane-sulfonimide (LiTFSI) in pure EDA or 1,3-DAP, bybiasing the working electrode at 3V with respect to a standard referenceelectrode, for a duration ranging from 5-60 min.⁴ This will result inthe electrodeposition of either L-PEI or L-PPI on the biased electrode,leaving the rest of the electrodes intact. Planar electrodes will bedefined microlithographically. These electrodes will be grown byevaporating first a thin layer of Ti to improve adhesion on theSiO₂-covered Si wafer, followed by the deposition of a thick layer or Ptand an optional second thin layer of Ti to enable the adhesion of a SiNoverlayer. This SiN layer is used to protect the edges of themicrolithographically defined Pt electrodes from delaminating in anaqueous environment. Following SiN patterning, the remaining Ti isstripped off by immersion of the wafer in a titanium etchant (i.e.,H₂SO₄/H₂O-1/1, 80° C.) to expose the underlying Pt layer.

Electroplating Ag on top of one of the patterned Pt electrode will beused to selectively grow the reference Ag/AgCl electrode. This can beaccomplished by electroplating in a solution comprising KCN, K₂CO₃, andKAg(CN)₂. Subsequent electrochemical oxidation of Ag to AgCl will takeplace at a constant current of 40 μA (at ˜0.5 V) in 0.1 M HCl forapproximately 10-30 minutes Since only the reference electrode isconnected to the voltage source, no deposition occurs on the otherelectrodes, which remain clean for the subsequent electrodeposition ofL-PEI or L-PPI (described above). The use of an auxiliary Pt electrodecan improve device reliability and long-term operation.

The fabrication of such a pH sensor is simple and straight-forward. Thethickness of the electropolymerized L-PEI or L-PPI are reported toinfluence the sensor response. In the case where there is interferenceof H₂O₂ with the pH sensor, this should to be quantified and included inthe multi-parameter sensor response characteristics.

In one embodiment, the sensor element comprises an electrochemicaloxygen sensor. Variations in the partial pressure of O₂ in the blood isexpected to have a significant effect on the glucose sensor response.This is because of the dual role of O₂ in GO_(x) enzymatic catalysis toform H₂O₂ and its subsequent oxidation to regenerate O₂. Providing anindependent assessment of O₂ concentration could improve our level ofconfidence in sensor accuracy and reliability. Design simplicity,stability and good current linearity over the range of oxygen from 0 to99.5% v/v have rendered electrochemical-based Clark sensors as thepreferred method for O₂ sensing. A number of planar miniaturizedversions of it have already been developed,⁴ and variations in these areoutlined below.

Planar electrodes will be defined microlithographically, as describedearlier. The Pt working electrode may be covered with a biocompatiblediffusion limiting membrane to control O₂ permeability. Fine tuning thethickness of this membrane aids in minimizing response time andmaintaining sensitivity.⁴ Layer-by-layer (LBL) growth of Nafion/Fe³⁺thin films allow for adjusting permeability of a variety of species. Byadjusting the pH, the conformation of film growth could be tailored soas to acquire films of desired thickness. A pH of 4.5, for example,induces surface spreading of Nafion onto the substrate, thus ensuring afilm growth consisting of surface spread and tightly meshed polymericchains that exhibit high tortuosity to permeation. Moreover, thepresence of Fe³⁺ groups prevents the potential calcification of thesefilms due to interactions of the negative SO³⁻ groups of Nafion and thephysiologically present Ca²⁺ions. This may be helpful to prevent in vivodegradation of these devices. The precise localization of such films maybe performed using the well-established technique of micro-contactprinting along with LBL assembly. Polyacrylamide hydrogels will beemployed for the construction of these stamps, defined by crosslinkingthem onto lithographically etched masters. The applied force of thehydrogel on the substrate and time of contact will be adjustedaccordingly.

The fabrication of this sensor is straightforward, although it requiresconsiderable skill in terms of integrating it with the other twoelectrochemical sensors on the same chip. Depending on feature size,stamp micropositioning is critical. Four-degree of movement (x, y, z andtilt) stages along with corresponding controllers may be helpful inmicropositioning.

In one embodiment, glucose sensor response is determined as a functionof temperature, pH and oxygen. As outlined above, the interdependence oftemperature, pH and oxygen content, together with various glucose levelsand film-specific construction parameters create a multi-dimensionalproblem. A system to integrate all variables into a single calibrationplatform would be useful.

Standardizing all basic elements of the sensors and keeping the numberof independent variables to a minimum is an objective after individualsensing functionality and longevity are established. This will befollowed by conducting a series of calibrations.

TABLE I Typical Glucose sensor voltages Sensor V_(REF) V_(Working) TypeCO₂ Sensor specific Sensor specific Electrochem Ionic Sensor specificSensor specific Electrochem Glucose 0.7 V 1.2 V Electrochem

1. An analyte sensing device comprising: an external control unit and animplantable sensor platform in wireless optical two-way operablecommunication, wherein the implantable sensor platform can pass though abore of a needle, wherein the implantable sensor platform comprises, inoperable communication, a photovoltaic device to receive optical powerfrom the external control unit to serve as a power source for poweringsaid implantable sensor platform, an optical receiver for detectingsignals produced by the external control unit, a plurality of sensorelements deposited on a surface of the implantable sensor platform andoperable for sensing one or more analytes, wherein the plurality ofsensor elements have one or more working electrodes, a referenceelectrode and a counter electrode in contact with the surface they aredeposited on such that the electrodes do not delaminate when exposed tobody fluids, an interfacing circuit, for providing operating parametersto the electrodes of the plurality of sensor elements and controlledfeedback for the operation of the plurality of sensor elements, whereinthe plurality of sensor elements generates a sensor output signal havinga sensor output signal magnitude proportional to the amount of analytepresent, wherein the interfacing circuit comprises at least onepotentiostat, a signal processing circuit interfaced with the sensoroutput signal, wherein the signal processing circuit converts the sensoroutput signal of the plurality of sensor elements to digital pulseshaving a pulse frequency, wherein the pulse frequency is determined bythe sensor output signal magnitude and wherein changes in the pulsefrequency are proportional to changes in the analyte levels, wherein thedigital pulses are transmitted to the external unit, a switching modeselector configured to cause the implantable sensor platform to performat least one of an initialization function, a power level checkfunction, a potentiostat circuit reconfiguration function for analytelevel measurement, an implantable sensor selection function, and animplantable sensor calibration function, one or more optical componentsfor facilitating wavelength selection, transmission and/or reflection,and a biocompatible coating surrounding at least a portion of theimplantable sensor platform, wherein said biocompatible coatingcomprises a drug and is designed to control the timed release of thedrug, wherein the external control unit comprises, in operablecommunication, an optical source suitable for powering the photovoltaicdevice of the implantable sensor platform, a receiver suitable forreceiving one or more digital pulses from the implantable sensorplatform and processing the one or more digital pulses to determine theanalyte levels, an optical transmitter suitable to transmit one or moreoptical pulses to the optical receiver of the implantable sensorplatform, wherein the optical pulse relays instructions to the switchingmode selector to cause the implantable sensor platform to perform atleast one of the initialization function, the power level checkfunction, the potentiostat circuit reconfiguration function, the sensorselection for analyte level measurement function, the implantable sensorselection function and the implantable sensor calibration function, anintegrated circuit for processing and displaying the analyte levels,wherein the integrated circuit is in operable communication with thereceiver, a microcontroller comprising a program code, programmablememory, and means to display output and communicate with other devices,means of interfacing with the receiver, optical source and opticaltransmitter, to establish an operable communication with the implantablesensor platform, a power supply to power the external unit, one or moreoptical components providing wavelength selection, transmission orreflection functions, and a miniaturized camera to align the implantablesensor platform with the optical components of the external controlunit.
 2. The analyte sensing device of claim 1, wherein the implantablesensor platform comprises a first sub-chip, a second sub-chip and athird sub-chip in operable communication, wherein the first sub-chipincludes, the photovoltaic device, a transmitter configured to transmitinformation in the form of digital pulses to the external unit, and theoptical receiver configured to receive instructions from the externalunit; the second sub-chip includes, the interfacing circuits includinginitialization, sensor select, and sensor calibration, the potentiostatand signal processing circuits, and the transmitter configured towirelessly transmit digital pulses; and the third sub-chip includes theplurality of sensor elements.
 3. The analyte sensing device of claim 2,wherein the first sub-chip, second sub-chip and third sub-chip areintegrated via at least one of through-Silicon-vias,partial-Silicon-vias and interconnects.
 4. The analyte sensing device ofclaim 1, wherein the interfacing circuit includes a voltage controllogic unit and the signal processing circuit includes a potentiostat andan analog to digital converter.
 5. The analyte sensing device claim 1,wherein the plurality of sensor elements are configured to monitor atleast one of the plurality of analytes.
 6. The analyte sensing device ofclaim 5, further comprising sensor-select and potentiostat circuits thatsequentially address the plurality of sensor elements.
 7. The analytesensing device of claim 1, further comprising a digital-to-RF convertercircuit and an antenna, wherein the digital-to-RF converter circuit isconfigured to receive and convert the digital pulses from the output ofthe signal processing unit into a wireless radio frequency(RF) signalresponsive to the analytes, and wherein the antenna is configured totransmit the RF signal to the external unit.
 8. The analyte sensingdevice of claim 7, wherein the external unit includes an external RFreceiver configured to receive the RF signal.
 9. The analyte sensingdevice of claim 1, further comprising a transducer configured to receiveand convert the digital pulses at the output of signal processing unitinto a wireless ultrasound signal responsive to the analytes.
 10. Theanalyte sensing device of claim 9, wherein the external unit includes anexternal ultrasound receiver configured to receive the wirelessultrasound signal.
 11. The analyte sensing device of claim 1 wherein theminiaturized camera is used to implant the implantable sensor platformunder the skin using a needle based insertion device.
 12. An analytesensing device comprising: an external control unit and an implantablesensor platform in wireless optical two-way operable communication,wherein the implantable sensor platform can pass though a 14 gauge orsmaller bore needle, wherein the implantable sensor platform comprisessub-chip #1, sub-chip #2 and sub-chip #3 and a biocompatible coatingsurrounding at least a portion of the sensor platform, wherein sub-chip#1 comprises: a photovoltaic device that powers the sub-chip #1,sub-chip #2 and sub-chip #3, a first optical receiver to receiveinstructions from a mode select unit located in the external controlunit via an external unit optical transmitter, wherein the external unitoptical transmitter is located in the external control unit, and asecond optical receiver for providing information regarding lightintensity received from light-emitting diodes located in the externalcontrol unit; wherein the first and second optical receivers operate atwavelengths such that the first and second optical receivers do notinterfere with each other, a sensor platform transmitter configured totransmit digital pulses, relaying information selected from sensoroutput, calibration, potentiostat check, or solar power level checkreceived from a driver located on subchip #2, one or more coatingsproviding wavelength selection, transmission or reflection functions,wherein sub-chip #2 comprises: a plurality of interfacing circuitsselected from initialization circuits, sensor select circuits, andsensor calibration circuits, a potentiostat, a signal processingcircuit, a logic circuit, a demultiplexer, a multiplexer, and the driverto enable transmission of feedback signals selected from a level ofradiation intensity received by the photovoltaic device, a referencevoltage of the potentiostat, or a sensor reading, and wherein the driveron subchip #2 transmits a plurality of digital pulses to the sensorplatform transmitter located on subchip #1 and transmits digital pulsesusing transmitter selected from wireless RF and ultrasound, whereinsub-chip #3 comprises a plurality of sensor elements operable forsensing of one or more analytes, and wherein the plurality of sensorelements has one or more working electrodes, a reference electrode and acounter electrode, wherein the one or more working electrodes, referenceelectrode and counter electrodes are in contact with the surface theyare deposited on in a way that they do not delaminate when exposed tobody fluids, wherein the plurality of sensor elements are in contactwith the potentiostat and other circuits located on subchip #2, whereinsub-chip #1, sub-chip #2 and sub-chip #3 are electrically interconnectedusing through-silicon-vias or partial-silicon-vias and interconnects andintegrated in a manner to operate in the presence of body fluids;wherein the external control unit comprises, in operable communication,an optical source comprising light-emitting diodes and laser diodes,wherein the optical source powers the photovoltaic device on sub-chip#1, a third receiver for receiving digital pulses from the sensorplatform transmitter on subchip #1 of the implantable sensor platform,wherein the signal processing circuit converts a sensor element outputsignal to digital pulses, wherein the sensor platform transmitterconverts the digital pulses to RF and ultrasound pulses, wherein thefrequency of the digital pulses is determined by the sensor elementoutput signal which is controlled by an analyte level in the bodyfluids, wherein the external unit optical transmitter transmits one ormore optical pulses to the first optical receiver, wherein the one ormore optical pulse relays instructions to sub-chip #2 for the switching,multiplexing, demultiplexing and logic circuits of sub-chip #2 toprovide at least one function wherein the at least one function includesan initialization function, a power check function, a potentiostatcircuit reconfiguration for analyte level measurement function, animplantable sensor selection function, and an implantable sensorcalibration function, an integrated circuit for processing anddisplaying an electrical pulse, wherein the integrated circuit is inoperable communication with an optical receiver, a microcontrollercomprising a program code; programmable memory; a means to display theoutput and communicate with other devices; a means of interfacing withthe optical source, a receiver located in the external control unitreceiving digital pulses from RF or ultrasound transducer serving astransmitter and an optical transmitter located in the external controlunit operating at 800-1000 nanometers, a power supply to power theexternal control unit, one or more optical components providingwavelength selection, transmission, or reflection functions, and aminiaturized camera to align the implantable sensor platform with theoptical components of the external control unit, wherein theminiaturized camera is used to assist in inserting the implantablesensor platform under the skin using a 14 gauge or smaller needle. 13.An implantable sensor platform delivery device for delivering animplantable sensor platform into the body of living being, comprising:an actuator, a plunger, and a needle cannula having a hollow cannulabore that extends the length of the cannula, wherein the hollow cannulabore is sized and shaped to moveably contain the implantable sensorplatform wherein the actuator is associated with the plunger and whereinthe plunger is configured to be moveably located within the hollowcannula bore, such that when the implantable sensor platform and theplunger is located within the hollow cannula bore, movement of theactuator causes movement of the plunger.